The present invention relates generally to implantable, radially expandable medical prostheses which are frequently referred to as stents. In particular, the present invention is a bioabsorbable self-expanding stent.
Self-expanding medical prostheses frequently referred to as stents are well known and commercially available. They are, for example, disclosed generally in the Wallsten U.S. Pat. No. 4,655,7.1, the Wallsten et al. U.S. Pat. No. 5,061,275 and in Hachtmann et al., U.S. Pat. No. 5,645,559. Devices are used within body vessels of humans for a variety of medical applications. Examples include intravascular stents for treating stenoses, stents for maintaining openings in the urinary, biliary, tracheobronchial, esophageal, and renal tracts, and vena cava filters.
A delivery device which retains the stent in its compressed state is used to deliver the stent to a treatment site through vessels in the body. The flexible nature and reduced radius of the compressed stent enables it to be delivered through relatively small and curved vessels. In percutaneous transluminal angioplasty, an implantable endoprosthesis is introduced through a small percutaneous puncture site, airway, or port and is passed through various body vessels to the treatment site. After the stent is positioned at the treatment site, the delivery device is actuated to release the stent, thereby allowing the stent to self-expand within the body vessel. The delivery device is then detached from the stent and removed from the patient. The stent remains in the vessel at the treatment site as an implant.
Stents must exhibit a relatively high degree of biocompatibility since they are implanted in the body. An endoprosthesis may be delivered into a body lumen on or within a surgical delivery system such as delivery devices shown in U.S. Pat. Nos. 4,954,126 and 5,026,377. Preferred delivery devices for the present invention include U.S. Pat. Nos. 4,954,126; 5,026,377. Suitable materials for use with such delivery devices are described in U.S. patent application Ser. No. 08/833,639, filed Apr. 8, 1997.
Commonly used materials for known stent filaments include Elgiloy(copyright) and Phynox(copyright) metal spring alloys. Other metallic materials than can be used for self-expanding stent filaments are 316 stainless steel, MP35N alloy, and superelastic Nitinol nickel-titanium. Another self-expanding stent, available from Schneider (USA) Inc. of Minneapolis, Minn., has a radiopaque clad composite structure such as shown in U.S. Pat. No. 5,630,840 to Mayer. Self-expanding stents can be made of a Titanium Alloy as described in U.S. patent application Ser. No. 08/598,751, filed Feb. 8, 1996.
The strength and modulus of elasticity of the filaments forming the stents are also important characteristics. Elgiloy(copyright), Phynox(copyright), MP35N and stainless steel are all high strength and high modulus metals. Nitinol has relatively low strength and modulus.
The implantation of an intraluminal stent will preferably cause a generally reduced amount of acute and chronic trauma to the luminal wall while performing its function. A stent that applies a gentle radial force against the wall and that is compliant and flexible with lumen movements is preferred for use in diseased, weakened, or brittle lumens. The stent will preferably be capable of withstanding radially occlusive pressure from tumors, plaque, and luminal recoil and remodeling.
There remains a continuing need for self-expanding stents with particular characteristics for use in various medical indications. Stents are needed for implantation in an ever growing list of vessels in the body. Different physiological environments are encountered and it is recognized that there is no universally acceptable set of stent characteristics.
A need exists for a stent which has self expanding characteristics, but which is bioabsorbable. A surgical implant such as a stent endoprosthesis must be made of a non-toxic, biocompatible material in order to minimize the foreign-body response of the host tissue. The implant must also have sufficient structural strength, biostability, size, and durability to withstand the conditions and confinement in a body lumen.
All documents cited herein, including the foregoing, are incorporated herein by reference in their entireties for all purposes.
The present invention is an improved implantable medical device comprised of a tubular, radially compressible, axially flexible and radially self-expandable structure including elongate filaments formed in a braid-like configuration. The filaments consist of a bioabsorbable polymer which exhibits a relatively high degree of biocompatibility.
Briefly, self-expanding stents of the present invention are formed from a number of resilient filaments which are helically wound and interwoven in a braided configuration. The stents assume a substantially tubular form in their unloaded or expanded state when they are not subjected to external forces. When subjected to inwardly directed radial forces the stents are forced into a reduced-radius and extended-length loaded or compressed state. The stents are generally characterized by a longitudinal shortening upon radial expansion.
In one preferred embodiment, the device is a stent which substantially consists of a plurality of elongate polylactide bioabsorbable polymer filaments, helically wound and interwoven in a braided configuration to form a tube. Bioabsorbable implantable endoprostheses such as stents, stent-grafts, grafts, filters, occlusive devices, and valves may be made of poly(alpha-hydroxy acid) such as poly-L-lactide (PLLA), poly-D-lactide (PDLA), polyglycolide (PGA), polydioxanone, polycaprolactone, polygluconate, polylactic acid-polyethylene oxide copolymers, modified cellulose, collagen, poly(hydroxybutyrate), polyanhydride, polyphosphoester, poly(amino acids), or related copolymers materials, each of which have a characteristic degradation rate in the body. For example, PGA and polydioxanone are relatively fast-bioabsorbing materials (weeks to months) and PLA and polycaprolactone are a relatively slow-bioabsorbing material (months to years).
A stent constructed of a bioabsorbable polymer provides certain advantages relative to metal stents such as natural decomposition into non-toxic chemical species over a period of time. Also, bioabsorbable polymeric stents may be manufactured at relatively low manufacturing costs since vacuum heat treatment and chemical cleaning commonly used in metal stent manufacturing are not required.
The present invention includes a method of designing and manufacturing an improved braided bioabsorbable stent which is different from practices used to make braided metal wire stents. The method involves selecting a specific bioabsorbable polymer based on a desired stent functional absorption time and stent radial force. The stent functional absorption time is the time period within which the stent retains at least 80% of its original radial strength. The stent is made by first selecting a braid design from the invention and making two different annealed stents. Radial force and dimensional test results from the two stents are used to develop a nearly linear mathematical equation to determine the parameters to meet the design goals. This method advantageously limits costly and time consuming trial and error to arrive at the optimum design.
Bioabsorbable polymer stents are radiolucent and the mechanical properties of the polymers are generally lower than structural metal alloys. Bioabsorbable stents may require radiopaque markers and may have a larger profile on a delivery catheter and in a body lumen to compensate for the lower material properties.
Bioabsorbable PLLA and PGA material are degraded in vivo through hydrolytic chain scission to lactic acid and glycolic acid, respectively, which in turn is converted to CO2 and then eliminated from the body by respiration. Heterogeneous degradation of semicrystalline polymers occurs due to the fact that such materials have amorphous and crystalline regions. Degradation occurs more rapidly at amorphous regions than at crystalline regions. This results in the product decreasing in strength faster than it decreases in mass. Totally amorphous, cross-linked polyesters show a more linear decrease in strength with mass over time as compared to a material with crystalline and amorphous regions. Degradation time may be affected by variations in chemical composition and polymer chain structures, and material processing.
PLA monofilaments may be produced by a process involving seven general steps as summarized herein. First, a polymer formed of poly-L-lactic acid is brought to an elevated temperature above the melting point, preferably 210xc2x0-230xc2x0 C. Second, the material is then extruded at the elevated temperature into a continuous fiber, by a conventional process, at a rate about of three to four feet per minute. Third, the continuous fiber is then cooled to cause nucleation. The cooling is preferably performed by passing the fiber through a nucleation bath of water. Fourth, the material then passes through a first puller, which runs at about the same speed as the extruder, and places the material under slight tension. Fifth, the fiber is then heated to a temperature between about 60xc2x0 C. and about 90xc2x0 C. (preferably 70xc2x0 C.) as it passes through a heated oven. To perform annealing, the oven can be designed to be quite long and heated near the end, so that the orientation and annealing take place in the same oven. Alternatively, a separate oven can be placed directly after the orientation oven. The annealing step heats the fibers to a range of about 65xc2x0 C. to about 90xc2x0 C., preferably closer to 90xc2x0 C. Sixth, while being heated in the orientation oven and the annealing oven, the fiber is drawn between the first puller located before the orientation oven and a second puller located after the annealing oven (if a separate oven). The material is drawn at a draw ratio of between about 5 to about 9, preferably between about 6 and about 8. Draw ratio describes the extension in length resulting from polymer extrusion or drawing. Quantitatively, the drawing ratio is a unitless value equal to the extruded or drawn length divided by the original length. Maintaining tension through the annealing step prevents shrinkage in later use. The second puller, located at the exit of the oven, runs at an increased speed necessary to provide the desired draw ratio. As the fiber exits the oven and passes through the second puller the tension is immediately released before the material cools. Seventh, finally, the fiber is collected onto spools of desired lengths.
Strength of the filaments generally increases with draw ratio and with lower draw temperatures. A draw ratio of between 5 and 9 is preferred. PLA is generally amorphous because of the material""s slow crystallization kinetics. Very slow cooling after drawing of the filament or use of a nucleating agent will cause crystallization. However, the material may be annealed at temperatures above about 60xc2x0 C. to cause crystallization, and generally, the strength decreases slightly and the modulus increases. Annealing is preferably performed after drawing to release residual stresses and to homogenize the surface to center variations in structure. Annealing will preferably be performed at a temperature of between about 60xc2x0 C. and 150xc2x0 C. for a period of time between about 5 and 120 minutes. Reference is made to Enhancement of the Mechanical properties of polylactides by solid-state extrusion, W. Weiler and S. Gogolewski, Biomaterials 1996, Vol 17 No. 5, pp. 529-535; and Deformnation Characterstics of a Bioabsorbable Intravascular Stent, Investigative Radiology, December 1992, C. Mauli, Agrawal, Ph.d., P.E., H. G. Clark, Ph.D., pp. 1020-1024. It is generally preferred in accordance with this invention that the annealed bioabsorbable filament has a substantially homogeneous cross-section, in other words, that it has a substantially solid cross-section without substantial variations between the center and the surface of the filament.
Mechanical properties generally increase with increasing molecular weight. For instance, the strength and modulus of PLA generally increases with increasing molecular weight. Degradation time generally decreases with decreasing initial molecular weight (i.e., a stent made of a low molecular weight polymer would be bioabsorbed before a stent made of a high molecular weight polymer). Low molecular weight PLA is generally more susceptible to thermo-oxidative degradation than high molecular weight grades, so an optimum molecular weight range should be selected to balance properties, degradation time, and stability. The molecular weight and mechanical properties of the material generally decreases as degradation progresses. PLA generally has a degradation time greater than 1 year. Ethylene oxide sterilization process (EtO) is a preferred method of sterilization. PLA has a glass transition temperature of about 60xc2x0 C., so care must be taken not to store products in environments where high temperature exposure may result in dimensional distortion.
PLA, PLLA, PDLA and PGA include tensile strengths of from about 40 thousands of pounds per square inch (ksi) to about 120 ksi; a tensile strength of 80 ksi is typical; and a preferred tensile strength of from about 60 ksi to about 120 ksi. Polydioxanone, polycaprolactone, and polygluconate include tensile strengths of from about 15 ksi to about 60 ksi; a tensile strength of 35 ksi is typical; and a preferred tensile strength of from about 25 ksi to about 45 ksi.
PLA, PLLA, PDLA and PGA include tensile modulus of from about 400,000 pounds per square inch (psi) to about 2,000,000 psi; a tensile modulus of 900,000 psi is typical; and a preferred tensile modulus of from about 700,000 psi to about 1,200,000 psi. Polydioxanone, polycaprolactone, and polygluconate include tensile modulus of from about 200,000 psi to about 700,000 psi; a tensile modulus of 450,000 psi is typical; and a preferred tensile modulus of from about 350,000 psi to about 550,000 psi.
PLLA filament has a much lower tensile strength and tensile modulus than, for example Elgiloy(copyright) metal alloy wire which may be used to make braided stents. The tensile strength of PLLA is about 22% of the tensile strength of Elgiloy(copyright). The tensile modulus of PLLA is about 3% of the tensile modulus of Elgiloy(copyright). Stent mechanical properties and self-expansion are directly proportional to tensile modulus of the material. As a result, a PLLA filament braided stent made to the same design as the metal stent has low mechanical properties and would not be functional. The invention advantageously provides polymeric braided stents with radial strength similar to metal stents and the required mechanical properties capable of bracing open endoluminal strictures.
A bioabsorbable PLLA braided tubular stent changes size when constrained onto a catheter delivery system and when deployed. A deployed PLLA stent is generally longer in length and smaller in diameter than a PLLA stent prior to loading. For example, PLLA stents that were initially 30 mm long with external diameters of about 10.7 mm had deployed lengths of about 90 mm with diameters of about 6.3 mm.
In comparison, a metal self-expanding stent generally has about the same dimensions before loading and after deployment. For metal stents, if it is known that the patient has a 9 mm diameter vessel, then a 10 mm metal stent (stent is intentionally oversized by about 1 mm) is loaded onto the delivery system for implantation. This rule is not applicable for a polymer stent because more oversizing is necessary.
The present invention provides improved polymeric stents and a method for designing and producing the improved polymeric stents whereby a polymeric stent of a certain size may be produced, loaded on the delivery system, and upon deployment will yield desired implant dimensions and have desired mechanical properties.
The present invention advantageously provides a bioabsorbable PLLA braided stent of a desired implant size, and provides a method to make the stent at a particular diameter (A), anneal the stent at a smaller diameter (B), and deploy the stent from a delivery system of diameter (C) whereby the stent will be xe2x80x9cprogrammedxe2x80x9d to self-expand to a desired implant diameter (D). The relationship between the diameters is A greater than B greater than D greater than C.
In sum, the invention relates to a bioabsorbable implantable stent having a tubular, radially compressible and self-expandable braided and annealed structure including a first set of between 5 and 18 filaments, each of which extends in a helix configuration along a center line of the stent and having a first common direction of winding. A second set of filaments of the same number as the first set, each extend in a helix configuration along a center line of the stent and having a second common direction of winding. The second set of filaments cross the first set of filaments at an axially directed angle of between about 120 and about 150 degrees when in a first free radially expanded state after being annealed, but before being loaded on a delivery device so as to form a plurality of interstices between filaments. The term xe2x80x9cfree statexe2x80x9d is used when no externally applied forces are acting on the device, for example, when the device is resting on a table. Each filament includes PLLA, PDLA, PGA, or combinations thereof and have a substantially solid and substantially uniform cross-section, a tensile strength of from about 40 ksi to about 120 ksi, a tensile modulus of from about 400,000 psi to about 2,000,000 psi, and an average diameter of from about 0.15 mm to about 0.6 mm. The first set of filaments and second set of filaments act upon one another to create an outwardly directed radial force sufficient to implant the stent in a body vessel upon deployment from a delivery device. The stent may have a second free radially expanded state after being loaded and then released from a deployment device and the first and second sets of filaments cross at an axially directed angle of between about 80 and about 145 degrees when in the second free radially expanded state. The second sets of filaments may crisscross at an axially directed angle of between about 90 and about 100 degrees when in the second free radially expanded state, and a second free state diameter of from about 3 mm to about 6 mm. The axially directed angle may be between about 110 degrees and about 120 degrees when in the second free radially expanded state. The stent may have an outside diameter when in the second free radially expanded state and the stent exerts an outwardly directed radial force at one half of the outside diameter of from about 40 grams to about 300 grams. The stent may have an implanted state after being loaded, released from a deployment device into a body vessel, and then implanted in the body vessel, with the first and second sets of filaments crossing at an axially directed angle of between about 95 and 105 degrees when the stent is in the implanted state. The stent may be radially constrained to half of its free diameter and the radial force, RF, exerted by the device, in grams, as a function of annealed diameter, D, in mm, is about RF=xe2x88x9215D+491xc2x120. The stent may be annealed at a temperature of from about 60xc2x0 C. to about 180xc2x0 C. for a period of time of from about 5 minutes to about 120 minutes. The stent may be annealed at a temperature of from about 130xc2x0 C. to about 150xc2x0 C. for a period of time of from about 10 minutes to about 20 minutes. The braid may be annealed to yield a crossing angle of from about 130 degrees to about 150 degrees. The stent may be further disposed in a stent delivery device and the filaments have a crossing angle of from about 30 degrees to about 120 degrees. The stent may be deployed from a delivery system into a body lumen and the filaments have a crossing angle of from about 70 degrees to about 130 degrees. The stent may provide structural integrity to a body lumen for less than about 3 years. The stent may further include polydioxanone, polycaprolactone, polygluconate, polylactic acid-polyethylene oxide copolymers, modified cellulose, collagen, poly(hydroxybutyrate), polyanhydride, polyphosphoester, poly(amino acids) and combinations thereof. The filaments may be mono-filament or multi-filament. The stent may substantially degrade in vivo in from about 1 year to about 2 years. xe2x80x9cSubstantially degradexe2x80x9d means that the stent has lost at least 50% of its structural strength. It is preferable that the stent lose about 100% of its structural strength. The filaments may include polyglycolide and the stent may substantially degrades in vivo in a time of from about 3 months to about 1 year. The filaments may further include polygluconate, polydioxanone, or combinations thereof and the stent may substantially degrade in vivo in from about 1 week to about 3 months. The stent may have at least one end of diminishing diameter so as to function as a filter. The filaments may be substantially homogeneous in cross section and length. The filaments may have a tensile modulus of from about 400,000 psi to about 1,200,000 psi. The filaments may have a tensile modulus of from about 700,000 psi to about 1,200,000 psi. The stent may includes a plurality of the filaments helically wound and interwoven in a braided configuration to form a tube.
The invention also relates to a method of using an implantable endoprosthesis including: providing a tubular, radially compressible, axially flexible, and radially self-expandable braided and annealed structure. The structure including from about 10 to about 36 elongate filaments. The filament comprising PLLA, PDLA, PGA, and combinations thereof. Each filament having a substantially uniform cross-section, a tensile strength of from about 40 ksi to about 120 ksi, and a tensile modulus of from about 400,000 psi to about 2,000,000 psi. The filaments disposed at an angle of from about 130 degrees to about 150 degrees in a free state, each filament having an average diameter of from about 0.15 mm to about 0.6 mm, and the stent having a radial force at one-half diameter of from about 40 grams to about 300 grams. The annealed structure having a first diameter; disposing the structure into a delivery system at a second diameter smaller than the first diameter; inserting the delivery system and endoprosthesis in a body lumen; deploying the endoprosthesis from the delivery system into the body lumen to a third diameter smaller than the first; and allowing the endoprosthesis to self expand in the body lumen to a fourth diameter greater than the third diameter.
The invention also relates to a method for treating a site within a vessel of a patient, including: providing a biocompatible medical device including a tubular and axially flexible braid-like annealed structure at a first diameter which is radially self-expandable between a compressed state and an expanded state and which includes from about 10 to about 36 elongate filaments. The filaments include PLLA, PDLA, PGA, and combinations thereof. Each filament has a substantially uniform cross-section, a tensile strength of from about 40 ksi to about 120 ksi, and a tensile modulus of from about 400,000 psi to about 2,000,000 psi; Providing a delivery system with the medical device positioned on a portion of the delivery system in the compressed state at a second diameter smaller than the first diameter, Inserting the portion of the delivery system with the medical device into the patient""s vessel at a location spaced from the treatment site, and manipulating the delivery system to advance the medical device through the vessel, to the treatment site; Deploying the medical device from the delivery system. The medical device being deployed at a third diameter smaller than the original free diameter and allowing the medical device to self-expand within the vessel; and Removing the delivery system from the patient with the medical device remaining in the expanded state and supporting the vessel.
The invention also relates to a bioabsorbable implantable device made from the process including providing a plurality of elongate filaments including PLLA, PDLA, PGA, and combinations thereof; braiding the filaments on a first mandrel to form a tubular, radially compressible, axially flexible, and radially self-expandable device. The device having a first diameter of from about 2 mm to about 10 mm larger than the final implanted device diameter; and annealing the device on a second mandrel having a second diameter smaller than the first diameter. The second mandrel diameter adapted to be computed from a linear equation relating radial force to annealed stent diameter. The equation being derived from measured radial force and measured annealed stent diameter data from two stent prototypes made on two anneal mandrel diameters and deployed from a device delivery system. Each filament may have a substantially uniform cross-section, a tensile strength of from about 40 ksi to about 120 ksi, and a tensile modulus of from about 400,000 psi to about 2,000,000 psi. Annealing may cause the device to radially shrink.
The invention also relates to a method of manufacturing a stent including: providing from about 10 to about 36 filaments consisting essentially of poly (alpha-hydroxy acid). The filaments have an average diameter from about 0.15 mm to about 0.60 mm; braiding the filaments at a braid angle of from about 120 degrees to about 150 degrees on a braid mandrel of from about 3 mm to about 30 mm diameter ; removing the braid from the braid mandrel; disposing the braid on an annealing mandrel having an outer diameter of from about 0.2 mm to about 10 mm smaller than the braid mandrel diameter; annealing the braid at a temperature between about the polymer glass-transition temperature and the melting temperature for a time period between about 5 and about 120 minutes; and allowing the stent to cool.
Bioabsorbable polymer resins are commercially available. Bioabsorbable resins such as PLA, PLLA, PDLA, PGA and other bioabsorbable polymers are commercially available from several sources including PURAC America, Inc. of Lincolnshire, Ill.
Still other objects and advantages of the present invention and methods of construction of the same will become readily apparent to those skilled in the art from the following detailed description, wherein only the preferred embodiments are shown and described, simply by way of illustration of the best mode contemplated of carrying out the invention. As will be realized, the invention is capable of other and different embodiments and methods of construction, and its several details are capable of modification in various obvious respects, all without departing from the invention. Accordingly, the drawing and description are to be regarded as illustrative in nature, and not as restrictive.